The invention relates generally to imaging systems. In particular, the invention relates to a scintillator for use in an imaging system and a method of making the same.
Positron emission tomography (PET) is a medical imaging technique in which a radioactive substance is administered to a patient and then traced within the patient's body by means of an instrument that detects the decay of the radioactive isotope. In PET, a chemical tracer compound having a desired biological activity or affinity for a particular organ is labeled with a radioactive isotope that decays by emitting a positron. The emitted positron loses most of its kinetic energy after traveling only a few millimeters in a living tissue. It is then highly susceptible to interaction with an electron, an event that annihilates both particles. The mass of the two particles is converted into 1.02 million electron volts (1.02 MeV) of energy, divided equally between two 511 keV photons (gamma rays). The two photons are emitted simultaneously and travel in almost exactly opposite directions. The two photons penetrate the surrounding tissue, exit the patient's body, and are absorbed and recorded by photodetectors typically arranged in a circular array. Biological activity within an organ under investigation can be assessed by tracing the source of the radiation emitted from the patient's body to the photodetectors.
The value of PET as a clinical imaging technique is in large measure dependent upon the performance of the photodetectors. Each photodetector comprises a scintillator cell or pixel coupled to photomultiplier tubes. The scintillator cell produces light at these two points that is sensed by the photomultiplier tubes. The electrical signals from the photomultiplier tubes are processed to produce an image of the patient's organ. The two photons, generated from annihilation of the positron, strike the scintillator cell at two points separated by 180 degrees. In other words, approximate simultaneous interaction of the photons on the scintillator cell indicates the presence of a positron annihilation along the line joining the two points of interaction. By measuring the slight difference in arrival times of the two photons at the two points in scintillator cell, the position of positron can be calculated.
The limitations of this time difference measurement are highly dependent on the stopping power, light output, and decay time of the scintillator material. Stopping power is the ability to stop the 511 keV photons in as little material as possible so as to reduce the overall size of the photodetectors and, therefore, enhance the light collection efficiency and energy resolution. The stopping power is typically expressed as the linear attenuation coefficient τ having units of inverse centimeters (cm−1). After a photon beam has traveled a distance x in a scintillator material, the proportion of photons that has not been stopped by the scintillator material is exp(−τ·x). Thus, for a good scintillator material, τ should be as large as possible. High light output is important because the photodetectors will have higher sensitivity and, thus, the dose of the radioactive material administered to the patient can be reduced. Decay time (or also known as time constant, decay constant, or primary speed) is a measure of how fast the scintillator material stops emitting light after a cessation of excitation by the 511 keV photon. Short decay time allows for more rapid scanning and, thus, better observation of the motion of the body's organs.
Known scintillator materials for PET are bismuth germanate (Bi4Ge3O12 or “BGO”), Lutetium orthosilicate (LSO) and thallium-doped sodium iodide (NaI:T1), NaI:T1 has a reasonable stopping power but a long decay constant of about 250 nsec (nanoseconds). BGO has a relatively good stopping power but a relatively low light output and a long decay constant of about 300 nsec.
Although lutetium silicate (LSO) offers good light yield and decay times, these scintillator materials are costly and difficult to fabricate. As a result, they are not economically viable for large area detector assemblies for a total body scanner using these scintillators.
Accordingly, there is a need for a suitable scintillator that addresses some or all of the problems set forth above.